Method of motion correction in optical coherence tomography imaging

ABSTRACT

An image data set acquired by an optical coherence tomography (OCT) system is corrected for effects due to motion of the sample. A first set of A-scans is acquired within a time short enough to avoid any significant motion of the sample. A second more extensive set of A-scans is acquired over an overlapping region on the sample. Significant sample motion may occur during acquisition of the second set. A-scans from the first set are matched with A-scans from the second set, based on similarity between the longitudinal optical scattering profiles they contain. Such matched pairs of A-scans are likely to correspond to the same region in the sample. Comparison of the OCT scanner coordinates that produced each A-scan in a matching pair, in conjunction with any shift in the longitudinal scattering profiles between the pair of A-scans, reveals the displacement of the sample between acquisition of the first and second A-scans in the pair. Estimates of the sample displacement are used to correct the transverse and longitudinal coordinates of the A-scans in the second set, to form a motion-corrected OCT data set.

PRIORITY

This application is a continuation of U.S. patent application Ser. No. 13/357,097, filed Jan. 24, 2012, which in turn is a continuation of U.S. patent application Ser. No. 12/794,926 (now U.S. Pat. No. 8,115,935), filed Jun. 7, 2010, which in turn is a continuation of U.S. patent application Ser. No. 12/075,477 (now U.S. Pat. No. 7,755,769), filed Mar. 11, 2008, which in turn is a continuation of U.S. patent application Ser. No. 11/331,567, filed Jan. 13, 2006 (now U.S. Pat. No. 7,365,856), which in turn claims the benefit of the filing date under 35 U.S.C. §119(e) of U.S. Provisional Patent Application Ser. No. 60/645,637, filed Jan. 21, 2005, each of which are hereby incorporated herein by reference in their entirety.

TECHNICAL FIELD OF THE INVENTION

The present invention relates to data acquisition methods for imaging by optical coherence tomography (OCT). In particular, the invention is a method for determining patient motion occurring during the acquisition of large sets of data by OCT. The method described herein acquires a sparse set of OCT data in a sufficiently short time that patient motion during the acquisition is not objectionable. The sparse set of OCT data acts as a set of guideposts for determination of the locations on the sample, of the measurements comprising the full data set.

BACKGROUND OF THE INVENTION

Optical Coherence Tomography (OCT) is a technique for performing high-resolution cross-sectional imaging that can provide images of tissue structure on the micron scale in situ and in real time [Huang et al. (1991)]. OCT is a method of interferometry that determines the scattering profile of a sample along the OCT beam. Each scattering profile is called an axial scan, or A-scan. Cross-sectional images, and by extension 3D volumes, are built up from many A-scans, with the OCT beam moved to a set of transverse locations on the sample. Motion of the sample with respect to the OCT scanner will cause the actual locations measured on the sample to be arranged differently than the scan pattern in scanner coordinates, unless the motion is detected and the OCT beam placement corrected to track the motion.

In recent years, frequency domain OCT techniques have been applied to living samples [Nassif et al. (2004)]. The frequency domain techniques have significant advantages in speed and signal-to-noise ratio as compared to time domain OCT [Leitgeb, R. A., et al., (2003); de Boer, J. F. et al., (2003); Choma, M. A., et al. (2003)]. The greater speed of modern OCT systems allows the acquisition of larger data sets, including 3D volume images of human tissue.

In the case of ophthalmology, a typical patient can comfortably hold his eye open for a few seconds. OCT systems can advantageously use these few seconds to collect extensive images. During such an acquisition, motion of the patients head and natural shifts in the patient's fixation will distort the image. Tracking the motion of the eye to correct the placement of the OCT beam has proven useful [U.S. Pat. No. 6,736,508; Hammer, D. X., et al. (2005)]. There is also motion along the OCT beam, which is not detectable by the common designs of eye trackers, but which does distort the OCT image.

There is therefore a need for a method to correct the placement of OCT image data acquired on a moving sample. The correction could be applied to the mechanism scanning the OCT beam, to approximately follow the motion of the sample. Alternatively, the correction could be applied when images are built from the A-scans acquired in the presence of sample motion. The need is for a method to determine the motion, in three dimensions, of the sample during the acquisition of the A-scans. A method that does not require an additional optical system for eye tracking would have the advantages of simplicity and lower cost.

SUMMARY OF THE INVENTION

The present invention acquires, in addition to the set of A-scans comprising the desired image, a widely-spaced set of guidepost A-scans that can be recorded quickly enough to avoid objectionable motion of the sample. This method compares some of the A-scans comprising the image with some of the guidepost A-scans. When comparison shows that the optical scattering profile of an image A-scan and guidepost A-scan closely match, the location of the image A-scan on the moving sample is assumed to be the same as that of the matching guidepost A-scan.

If one were to assume no sample motion, one would expect the matches to be found during the course of the imaging scan pattern, whenever the OCT beam probes the same location, with respect to the scanner, as it did when collecting one of the guideposts (i.e. the scanner coordinates of the OCT system would be the same). Motion of the sample will cause the matches to occur for A-scans recorded at somewhat different scanner coordinates than would have been expected under the assumption of no sample motion. Each time a match is found, comparison of the actual scanner coordinates and the expected scanner coordinates (under the assumption of no sample motion), reveals the transverse displacement of the sample between acquisition of the guidepost scans and the acquisition of the matching A-scan in the image set. Comparison of the contents of the matching pair of A-scans reveals any longitudinal displacement of the sample, which would appear as a longitudinal shift in the image data between the matching pair of A-scans.

The comparison between A-scans need only be done between pairs of A-scans that are likely to match, such as those pairs which would have measured nearly the same location on the sample in the absence of sample motion. Depending on the comparison method, many pairs may match to some degree, so the method chooses the best-matching pair.

During the course of the imaging scan pattern, each match found provides updated information on the displacement of the sample. One can estimate the position of the sample between such matches by fitting a smooth curve through the points determined by the matches. Given the resulting curve of sample position versus time, the image data can be shifted to its correct locations, to form a 3D image free of motion artifact.

Other methods of determining eye motion use a landmark, such as the optic disk. The landmark is identified first, and its location is monitored as the detailed OCT scan proceeds. The landmark can be tracked on a separate imaging system, or the OCT beam can scan the landmark occasionally, briefly interrupting the larger data acquisition. However, good landmarks are not always found in diseased tissue. The method disclosed here takes advantage of the fact that the details in the structure of any tissue can serve the same need as a landmark.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates schematically one type of OCT system.

FIG. 2 is an example of an OCT image of human eye tissue, showing the detail that exists within the image.

FIG. 3 shows three views of a 3D OCT data set collected from a human eye, in which motion artifact is visible.

FIG. 4 shows three views of a 3D OCT data set collected from a human eye, in which motion artifact has been largely corrected by the methods disclosed herein.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT

FIG. 1 is a schematic illustration of one design of OCT system. The light source 101 provides light having a short-coherence length to the fiber-based interferometer. Directional coupler 102 serves to split the light from source 101 between a reference arm 103 and a sample arm 104. Lens 111 and mirror 112 serve to return reference light back to coupler 102. A scanning system including lenses 121 a, 121 b, and 121 c, and scanning mirrors 122, directs light successively along paths such as 123 a, 123 b, 123 c onto successive locations on sample 125. Some light scattered from sample 125 returns closely enough along the illumination path to re-enter the fiber interferometer, is combined with reference light in coupler 102. The interfered sample and reference light are detected in detector 130.

The scanning mirror 122 is controlled by a system processor and generates scan coordinates which correspond to certain transverse positions on the sample. A sample, such as the human eye, will move with respect to the OCT system. Once the sample moves the scanner coordinates associated with a particular transverse position on the sample will change in an unknown manner. The subject application describes a method for determining the extent of this displacement and correcting for that displacement.

As noted above, the sample 125 may move with respect to the measurement system causing a time-dependent difference between scanner coordinates and sample coordinates. In some OCT systems, such as handheld scanners, motion of the scanning optics can contribute to the relative motion between scanner coordinates and sample coordinates.

FIG. 2 shows an example OCT tomogram from human tissue. Each column in this image is one A-scan; successive A-scans at neighboring locations in tissue are placed side-by-side to form the image. The grey scale is inverted, with highly scattering regions shown dark, to better reproduce the structure within the image that this method uses to recognize regions of tissue.

Referring to FIG. 3, a three dimensional data set acquired from a human eye is shown in three views. En-face projection 300 is the integrated optical scattering, integrated along depth z to produce a view similar to what one would see in an ophthalmoscope. Sections 301 and 302 show sections of the volume in which one axis z is the direction of propagation of the OCT beam. Section indicators 301 a and 302 a show the locations of sections 301 and 302, respectively, in projection 300. Sections such as 301 and 302 are called tomograms. The columns comprising tomogram 301, and the rows comprising tomogram 302, are individual A-scans. Each pixel in the projection 300 is the integrated intensity from one A-scan. The dimensions of the tissue within the image are approximately 3 mm along each of x and y, 1.5 mm along z.

The acquisition of A-scans for this example proceeded along the horizontal rows of projection 300. Tomogram 301 is one of these rows, so the A-scans comprising this section were acquired sequentially. Tomogram 302 consists of A-scans chosen from successive rows, so one sees in section 302 the motion of the sample during the duration of the acquisition of the full 3D volume.

In addition to the raster of A-scans, we acquire a set of guidepost A-scans, which is represented in FIG. 3 by the lower-resolution en-face image 350 composed of the integrated optical scattering in the set of guidepost A-scans en-face image. The preferred embodiment uses approximately a 32-by-32 grid of guidepost A-scans, which can be acquired within 100 milliseconds by modern OCT techniques. For purposes of illustration, the example of FIG. 3 uses a guidepost set consisting of 30 rows of 128 A-scans; this larger set produces a more recognizable en-face image 350, while still requiring less than 200 milliseconds for acquisition by modern OCT techniques.

The set of guidepost A-scans is preferably acquired quickly enough that the sample is substantially stationary, meaning that there is no objectionable motion of the sample. For example, when imaging the retina of the human eye, the transverse optical resolution is typically no better than 5 microns. The fast motions of the eye called tremor cause only a few microns apparent motion of the retina, apparent motions being the motion as seen through the optics of the human eye. Motions due to tremor are typically not resolvable, so they are not considered objectionable. The jerk-like eye rotations called saccades can cause 100 microns apparent motion, so they are objectionable. Saccades occur irregularly, approximately once per second. Thus when applied to OCT of the human eye, the set of guidepost A-scans must be acquired within significantly less than one second in order to avoid objectionable motion. Acquisition of the guidepost A-scans within 200 milliseconds is preferred in this application.

For comparison purposes, the full raster scan of image information shown in FIG. 3 corresponds to 256 rows of 256 A-scans. Using current technology, this raster scan can take on the order of 1 to 2 seconds to perform during which time the eye will move.

There are alternatives to a raster scan for collecting the OCT imaging information. One can use a set of radial scans, scanning transversely across a set of lines that extend outward from a center point; radial scans in this pattern have the characteristic of higher density sampling near the crossing point. Another alternative pattern is a set of concentric circles. Because the currently described method takes advantage of revisiting the portions of the sample seen during collection of a set of guideposts, the method provides the greatest advantage when the scan pattern for the full data set covers a two-dimensional area. Scan patterns with a two-dimensional transverse extent can be collected using a scan pattern designed to frequently cross a set of guideposts, with the guideposts conveniently collected using one or a few transverse scan lines. A two-dimensional extent in this context is an area having transverse dimensions more than ten times the transverse optical resolution, so that meaningful tomogram images can be extracted in each transverse dimension, such as tomograms 301 and 302 in FIG. 3.

The guidepost A-scans are compared with A-scans in the raster for the purpose of finding matches of the scattering profile. One effective comparison method is the normalized cross-correlation, defined by X ₁₂(t)≡∫a ₁*(z)·a ₂(z−t)dz/√{square root over (∫a ₁ ² dz·∫a ₂ ² dz)} where a(z) is a measure of the scattering intensity along the OCT beam, t is a variable indicating the relative shift along the beam direction of the two A-scans, and the * denotes complex conjugation. The A-scans a₁ and a₂ are chosen from the guidepost set and the raster, respectively.

If the A-scans were measured at locations more closely spaced than the width of the OCT beam, then there will likely be one A-scan in each row of the full data set which correlates well with one guidepost A-scan. If there has been eye motion in x and y between acquisition of the guideposts and acquisition of the raster scan, the pair of A-scans with the highest cross-correlation will consist of two A-scans acquired at different sets of scanner coordinates of the OCT system. The difference in scan coordinates reveals the transverse displacement of the sample between acquisition of the guideposts and acquisition of the matching A-scan in the raster.

The cross-correlation is a function of the shift t. The value of t at which X₁₂(t) is maximized reveals the longitudinal displacement of the sample, along z, between acquisition of the guideposts and acquisition of the matching A-scan in the raster. In this way the displacement in x, y and z of the sample is determined at several points along the raster scan, each time a good match is found between an A-scan in the raster and an A-scan in the set of guideposts. The location of the sample between such matches is estimated by fitting a smooth curve, as a function of progress time of the raster scan, through the points determined by the matches. Given the resulting curve of sample displacement versus time, the data in the raster of A-scans can be shifted and interpolated to form a three-dimensional data set that is free of motion artifact.

The method disclosed herein is most effective when used to accurately correct minor movements during acquisition of the image A-scans. It can be advantageously combined with prior art methods that are better suited to correct sudden motions, such as saccades of the eye. One prior art method that is better suited to correction of saccades is cross-correlation of each tomogram in the image set of A-scans with its immediate neighbors; this prior art method relies on continuity of anatomical features between tomograms. The method disclosed herein could also be advantageously combined with coarse tracking of sample motion; coarse tracking may be more practical to implement than tracking accurate enough to eliminate all objectionable displacement between the desired and actual measurement locations.

FIG. 4 shows the data of FIG. 3 after correction using the method described herein. The data contained in the image A-scans has been shifted, and interpolated as necessary, so that each image A-scans that matched a guidepost A-scan now appears in a position corresponding to the location where that guidepost A-scan was acquired. En-face projection 400 is built from the optical scattering in this corrected data set, integrated along depth z. Tomograms 401 and 402 are sections through the corrected data set, related to projection 400 as indicated by section indicators 401 a and 402 a respectively. Indication 450 shows the location extent of the guidepost set 350 in relation to the full set of image A-scans.

The motion artifact seen in tomogram 302 is corrected in tomogram 402. Tomogram 402 represents the true shape of the imaged tissue section. Correction of the en-face projection 400 is more subtle, but can be noted by comparing it with en-face projections 300 and 350.

An alternative pattern for the set of guideposts would be to form one or more generally diagonal sections across the volume to be covered by the full data set. Then in the absence of motion, each row of A-scans in the raster will be likely to match one of the A-scans in each diagonal section. However, there can be occasional erroneous matches between A-scans acquired at two different locations in tissue, when the tissue at these two locations happens by chance to have similar optical scattering profiles. Such erroneous matches would cause distortion in the output. The grid pattern of guidepost A-scans used in the examples of FIG. 3 and FIG. 4 tends to provide clusters of matches to the image A-scans; that is, some rows in the raster contain several neighboring A-scans that have matching guidepost A-scans. Having matches occur in clusters, coupled with the knowledge that matches within a cluster should show nearly the same sample displacement due to their proximity in time, plus a step of fitting the displacements determined by the matches to a smooth curve in time, gives protection against erroneous matches. Patterns of guideposts that cause matches to occur in clusters are therefore preferred.

A number of practical complications can be accommodated in this method. There are some choices for the function a(z) used to represent each A-scan. In many methods of OCT, including the modern Fourier-domain methods, the information on the amplitude and phase of the scattered light is recorded; a(z) could be chosen to be this complex-valued function. However, between acquisition of the guidepost scans and acquisition of the corresponding A-scans in the raster, the relative phases of the light scattered from different depths may have changed. This will happen if scattering centers move relative to one another by even a fraction of a wavelength of the probe light. Such motion is likely in living tissue, and such changes in a(z) would prevent one from finding a good match. For this reason, a(z) is best taken to be the real-valued intensity of light scattered from depth z, so that the comparison ignores the phase information.

OCT typically measures the scattering with an axial resolution of a few microns. Motion of some tissues, such as blood vessels, can change the scattering profile over this length scale. Also, if the spacing of A-scans in the raster is not much less than the OCT beam diameter, then no A-scan in the raster will exactly match any guidepost A-scan. One can apply smoothing of the scattered intensity, as measured by OCT, and use the values in the smoothed data set to form the function a(z). The preferred amount of smoothing retains as much detail in the A-scans as can be expected to be found upon re-measuring the same tissue, given the motion within the tissue and the coarseness of the raster scan.

Failures to find well-correlated A-scans will still occur. Eye motion during acquisition of a row in the raster could cause a failure, if the eye motion causes the raster to skip over the region covered by one of the diagonal traces in the guidepost set. Intermittent failures in the scan, due to dust on optics or floaters in the eye, could make a few a-scans un-useable. Eye motion greater than the extent of the region scanned will also cause registration failures. If the comparison method is cross-correlation, failed matches are recognized by their low cross-correlation coefficient, relative to the coefficient found in non-moving similar tissue. The effect of failed matches is mitigated by the step fitting of a smooth curve through the displacements determined by successful matches. Failed matches can be omitted from the fit, or the cross-correlation coefficient can be used as a confidence value in weighted fitting.

We consider now efficient application of the preferred embodiment to Fourier-domain methods of OCT, which record interference spectra at each transverse location on the sample. It is desirable to have a method of finding the proper position of an A-scan quickly directly from the spectrum, without taking the Fourier transform required to reconstruct the A-scan. Calculating the displacement of the sample quickly is desired to allow correction of the raster scan position during the raster scan, so that areas are not skipped due to motion of the patient against the direction of the scan.

The cross-correlation X₁₂ can be efficiently calculated from the spectra that are directly provided in Fourier-domain methods of OCT. The cross-correlation between two A-scans is the Fourier transform of the product of the two corresponding fringe spectra (the portion of the spectra due to interference, after subtracting the intensity of the reference beam). Each A-scan is related to a fringe spectrum through a(z)=∫e ^(i2qz) S(q)dq, where S(q) is the fringe spectrum as a function of the optical frequency q, with q in units of radians of optical phase per unit length. The cross-correlation is related to the fringe spectra by X ₁₂(t)≡∫e ^(i2q·t) S ₁*(q)·S ₂(q)dq/√{square root over (∫S ₁ ² dz·∫S ₂ ^(w) dz)}.

The cross-correlation between a pair of A-scans, if the A-scans are recorded at the same transverse location on the patient to within the beam diameter, has a peak at a position corresponding to the motion along the z-axis between acquisition of the two A-scans. The size of these peaks will be greatest for the pair of A-scans that best overlap in the transverse directions. Thus, the integrated squared magnitude of the cross-correlation indicates quantitatively the goodness of match between a pair of A-scans, and this integrated magnitude is the same as the integrated squared magnitude of the point-by-point product of the corresponding spectra, by Parseval's theorem. The integrated squared magnitude of the cross-correlation is efficiently calculated from C ₁₂ =∫|X ₁₂(z)|² dz≡∫|S ₁(q)S ₂*(q)|² dq/√{square root over (∫S ₁ ² dq·∫S ₂ ² dq)}.

The quantity C₁₂, computed for each pair of A-scans that is a candidate match, is used to identify matching pairs and thus the sample displacement in x and y. Then the value of the shift t at which X₁₂ has its peak value is used to determine the sample displacement in z.

One further complication is that Fourier-domain OCT techniques that record a single interference spectrum for each A-scan, with a single-phase of the reference beam, produce a double image. Each A-scan a(z) consists of the reflectivity profile along z, plus the mirror image of that profile. The cross-correlation is complicated by this mirror image, having two peaks instead of one: a peak at the position corresponding to the z-motion of the patient between acquisitions and another peak at the position corresponding to the negative of the z-motion of the patient.

One can avoid this problem by eliminating the part of the A-scan corresponding to positions z<0. Working directly on the spectra, one can eliminate this part of the corresponding A-scan by convoluting with a kernel that has a Fourier transform that is 1 for most of the z>0 axis, and 0 for most of the z<0 axis. An example of such a kernel is (−i/4, ½, +i/4). Convolution with a short kernel such as this one is more efficient than computing the Fourier transform; the convolution operation could be seen as an efficient approximation to the Hilbert transform. The resulting fringe spectra are complex valued and have a Fourier transform that is largely attenuated along the range z<0. After performing this convolution, the cross-correlation can be computed as above.

Thus an efficient method to obtain the integrated squared magnitude of the cross-correlation between two A-scans, using only the two spectra comprising the SD-OCT measurements of two A-scans, is as follows: 1) Isolate the fringe part of the spectra by subtracting the reference beam. 2) Perform an approximate Hilbert transform on each spectrum, for example by convolution with the truncated kernel described in the paragraph above. 3) Form the sums of the pointwise products of pairs of transformed spectra, to obtain cross-correlations C₁₂.

An extension of this fast cross-correlation technique provides an alternative method to estimate the sample displacement along z. The cross-correlation has a peak at a position corresponding to the motion along z, and the magnitude of the cross-correlation is symmetric about that peak. Therefore an estimate of the motion can be extracted by computing the average value of z, weighted by the squared magnitude of the cross-correlation:

$\begin{matrix} {{\Delta\; z} = {\frac{1}{C_{12}}{\int{{z \cdot {{X_{12}(z)}}^{2}}{\mathbb{d}z}}}}} \\ {= {\frac{1}{C_{12}}{\int{{S_{1}^{*}(q)}{S_{2}(q)}\frac{1}{2i}{\frac{\mathbb{d}}{\mathbb{d}q}\left\lbrack {{S_{1}(q)}{S_{2}^{*}(q)}} \right\rbrack}{{\mathbb{d}q}/\sqrt{\int{S_{1}^{2}{{\mathbb{d}q} \cdot {\int{S_{2}^{2}{\mathbb{d}q}}}}}}}}}}} \end{matrix}$ The longitudinal displacement is given by Δz calculated by the formula above, instead of by searching the function X₁₂(t) for its peak value.

This method allows several variations and extensions. The scan pattern collecting the main OCT image need not be a raster scan, but can be any set of A-scans that build a useful OCT data set. This method is useful so long as a set of guidepost A-scans can be acquired in locations that are likely to be included in the full OCT data set, given expected motion of the sample. The set of A-scans serving as guideposts can be acquired before or after the main OCT image. The comparison function need not be a cross-correlation, but can by any measure of the likelihood that two A-scans have measured the same location on the sample. Some alternate methods to compare candidate pairs of A-scans are: their mean-square difference, their Kullback Leibler distance, or the mutual information metric between the candidate pair. The comparison method can be applied to OCT optical scattering data in which the phase of the scattering is preserved, to OCT optical scattering data reduced to an intensity, or to logarithmically scaled intensities. Methods to reduce random noise in OCT data, such as thresholding and median filtering, have been described in the art and can be applied to the OCT data before comparison.

Although various embodiments that incorporate the teachings of the present invention have been shown and described in detail herein, those skilled in the art can readily devise many other varied embodiments that still incorporate these teachings.

The following references are hereby incorporated herein by reference.

US PATENT DOCUMENTS

-   -   Application US20050140984 Hitzenberger. “Efficient optical         coherence tomography (OCT) system and method for rapid imaging         in three dimensions”     -   U.S. Pat. No. 6,736,508 Xie, Kohnle, and Wei. “Tracking assisted         optical procedure”     -   U.S. Pat. No. 4,937,526 Ehman and Felmlee “Adaptive Method for         Reducing Motion and Flow Artifacts in NMR Images”     -   Co-pending U.S. patent application Ser. No. 10/750,341,         Hitzenberger, P. “Efficient optical coherence tomography (OCT)         system and method for rapid imaging in three dimensions”     -   US20030199769, Podoleanu, A. et al. (2002) “Apparatus for high         resolution imaging of moving organs” (WO03086181)

OTHER PUBLICATIONS

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What is claimed is:
 1. A method of imaging an eye with an optical coherence tomography (OCT) system, said system having a light source for generating a beam of radiation that is scanned over the eye with a scanner that generates scanner coordinates, said method comprising: scanning the beam in a raster pattern consisting of a plurality of parallel first lines, each first line including a plurality of A-scans; scanning the beam along at least one second line, said second line including a plurality of A-scans, said second line being oriented along an axis that is not parallel to the first lines, with certain of the A-scans of the second line being close to certain A-scans in the first lines; comparing the A-scans of the first and second lines that are near each other to determine any differences between the generated scanner coordinates and expected scanner coordinates, said differences being associated with eye movement; and using the determined differences to adjust image information derived from the A-scans to compensate for eye movement.
 2. A method as recited in claim 1 further including scanning the beam along a plurality of second lines, each of said second lines being oriented along an axis that is not parallel to the first lines.
 3. A method as recited in claim 2 wherein said second lines are diagonal to the first lines.
 4. A method as recited in claim 1, wherein the determined differences are used to determine longitudinal displacement.
 5. A method as recited in claim 1, wherein the determined differences are used to determine transverse displacement.
 6. A method as recited in claim 1, wherein the determined differences are used to determine both longitudinal and transverse displacements.
 7. A method as recited in claim 1, wherein the number of first scan lines is larger than the number of second scan lines.
 8. A method as recited in claim 7 wherein the second scan lines are acquired in a time period short enough so that the movement of the eye is less than the amount that would cause objectionable image distortion.
 9. A method as recited in claim 7 wherein the second scan lines are acquired in less than 200 milliseconds.
 10. A method as recited in claim 1, further comprising the step of generating an image using the adjusted image information.
 11. A method as recited in claim 1, wherein the scattered intensity of the A-scans in the first and/or second line is smoothed prior to the comparing step.
 12. A method as recited in claim 1, further including the step of tracking the motion of the eye during the scanning step.
 13. A method as recited in claim 1, wherein the comparing step includes taking cross-correlations between A-scans in the first and second lines.
 14. An optical coherence tomography (OCT) device comprising: a light source for generating a beam of radiation; a beam divider for directing a first portion of the light into a reference arm and a second portion of the light into a sample arm; optics for scanning the beam in the sample arm over a set of transverse locations on a sample; a detector for measuring light radiation returning from the sample and reference arms and generating output signals in response thereto; and a processor for converting the output signals into three dimensional image information; said processor for controlling the scanning optics in a manner to scan the beam in a raster pattern consisting of a plurality of parallel first lines, each first line including a plurality of A-scans and associated scanner coordinates; said processor for scanning the beam along at least one second line, said second line including a plurality of A-scans, said second line being oriented along an axis that is not parallel to the first lines, with certain of the A-scans of the second line being close to certain A-scans in the first lines; and wherein said processor compares the A-scans of the first and second lines that are near each other to determine any differences between the scanner coordinates associated with the respective A-scans and expected scanner coordinates, said differences being associated with eye movement and wherein said processor uses the determined differences to adjust image information derived from the A-scans to compensate for eye movement.
 15. A device as recited in claim 14 wherein said processor functions to scan the beam along a plurality of second lines, each of said second lines being oriented along an axis that is not parallel to the first lines.
 16. A device as recited in claim 15 wherein said second lines are diagonal to the first lines.
 17. A device as recited in claim 14, wherein the determined differences are used to determine longitudinal displacement.
 18. A device as recited in claim 14, wherein the determined differences are used to determine both longitudinal and transverse displacements.
 19. A device as recited in claim 14, wherein the number of first scan lines is larger than the number of second scan lines.
 20. A device as recited in claim 14, further comprising the step of generating an image using the adjusted image information.
 21. A device as recited in claim 14, wherein the comparing step includes taking cross-correlations between A-scans in the first and second lines. 